A wide variety of metal alloys are used to manufacture medical devices such as, but not limited to, guide wire devices, implantable vascular endoprostheses (e.g., stents), embolic protection filters, closure elements, and the like. Because of their high degree of biocompatibility and durability, nickel-titanium (“Ni—Ti”) alloys are used for fabricating many medical devices.
For example, guide wires are used to guide a catheter for treatment of intravascular sites such as PTCA (Percutaneous Transluminal Coronary Angioplasty), or in examination such as cardio-angiography. For example, a guide wire used in the PTCA is inserted into the vicinity of a target angiostenosis portion together with a balloon catheter, and is operated to guide the distal end portion of the balloon catheter to the target angiostenosis portion.
A guide wire needs appropriate flexibility, pushability and torque transmission performance for transmitting an operational force from the proximal end portion to the distal end, and kink resistance (resistance against sharp bending). To meet such requirements, Ni—Ti alloys and high strength materials (e.g., stainless steel) have been used for forming a core member (wire body) of a guide wire.
Near equi-atomic binary nickel-titanium alloys are known to exhibit “pseudo-elastic” behavior when given certain cold working processes or cold working and heat treatment processes following hot working. Pseudo-elasticity can be further divided into two subcategories: “non-linear” pseudo-elasticity and “linear” pseudo-elasticity. “Non-linear” pseudo-elasticity is sometimes used by those in the industry synonymously with “superelasticity.”
“Non-linear” pseudo-elastic Ni—Ti alloy exhibits upwards of 8% elastic strain (fully-recoverable deformation) by virtue of a reversible, isothermal stress-induced martensitic transformation. Non-linear pseudo-elasticity is known to occur due to a reversible phase transformation from austenite to martensite, the latter more precisely called “stress-induced martensite” (SIM). At room or body temperature and under minimal stress the material assumes a crystalline microstructure structure known as austenite. As the material is stressed, it remains in the austenitc state until it reaches a threshold of applied stress (a.k.a. the “upper plateau stress”), beyond which the material begins to transform into a different crystal structure known as martensite. Upon removal of the applied stress, the martensite reverts back to the original austenite structure with an accompanying return to essentially zero strain (i.e., the original shape is restored).
A “linear” pseudo-elastic Ni—Ti alloy is processed by cold working the material (e.g., by permanently deforming the material such as by wire-drawing) without subsequent heat treatment (i.e., partial or full annealing). Residual permanent deformation, i.e., “cold work,” tends to stabilize the martensitic structure so its reversion back to austenite is retarded or altogether blocked. With increasing levels of permanent deformation, the otherwise austenitic material becomes fully martensitic at room and body temperature, and further permanent deformation serves to progressively raise its yield strength. The complete disappearance of austenite via cold work altogether eliminates the plateau (austenite to martensite transformation) on the stress strain curve, and results in a unique stress strain curve without a classic perfectly linear slope and without an apparent yield point.
While linear binary NiTi is highly durable with good flexibility, binary NiTi may not be an ideal material for certain medical devices due to its inherently low stiffness [i.e., secant modulus around 5 Msi (˜34 GPa) at 4% elongation versus an elastic modulus of approximately 28 Msi (˜193 GPa) for 316L austenitic stainless steel]. For example, the low modulus of the material in the martensitic condition (either linear pseudo-elastic martensite or stress-induced martensite found in superelastic Ni—Ti) relative to an austenitic stainless steel makes it challenging to torque a guide wire made from linear pseudo-elastic Ni—Ti alloy because it has a greater tendency to elastically absorb a significant amount of applied twist as opposed to directly transmitting torque from end to end. Further, Ni—Ti has only moderate plateau stress levels, and is therefore less resistant to bending forces (as compared to stainless steel), and thus less effective at providing support as a guide wire for catheter delivery or as a stent for arterial scaffolding.